Conventional fluorescent activated cell sorters (FACS) are widely used in research and clinical applications1. These instruments are capable of very fast, multiparameter analysis and sorting but generally require large sample volumes, a trained operator for operation and maintenance, and are difficult to sterilize. FACS instruments are able to analyze as few as 10,000 and as many as tens of millions of cells. However, below 100,000 cells the ability to perform sorting diminishes1. Other separation methods such as magnetic beads don't require as many cells as FACS but they suffer from nonspecific binding, aggregation of cells and beads, and from the possibility that the beads themselves could interfere in subsequent processing steps. Thus, for sorting precious, small samples or cells from primary tissue, a cell sorter that is capable of handling small sample volumes with low cell numbers and that allows efficient recovery of the sorted populations addresses a unique scientific niche.
Microfabricated cytometers have the potential to sort with as few as 1,000 cells while concomitantly consuming less reagents in an easy to use, closed system. The latter is important because, unlike conventional FACS instruments, aerosols are not created, reducing the risks of contamination of the sorted cells and of working with biohazardous materials. Several microfabricated cell sorters have been described, but mostly as “proof of concept”. Fu, et al.2 reported 30-fold enrichment of E. coli at a throughput of 17 cells/s. Only 20% of the bacteria were viable after sorting and the sort purity in the target reservoir was 30%. In a subsequent study3, the throughput increased to 44 cells/s but the target purity decreased to less than 10%, with recovery reported as 39%. Wolff, et al.4 were able to sort beads from chicken red blood cells at a throughput of 12,000 events/s, with 100-fold enrichment. However, purity in the target well was about 1%. In these studies, enrichment was defined as the increase in the concentration of the target population in the collection well compared to the starting concentration. Purity referred to the accuracy of the sort and was the percentage of target cells sorted over all cells sorted into the collection well. Recovery was defined as the number of cells counted by the fluorescent detector vs. cells recovered from the collection well. The latter two studies used pressure switches in microfluidic devices that switched the entire fluid flow path and, consequently, any particles contained within the fluid plug. The mechanical compliance in these switches caused the fluid switch speed to be the rate limiting step for throughput3. Electrokinetic flow control has also been reported, e.g., electroosmosis2,5,6 or dielectrophoresis7,8,9, but the high electric field gradients and physicochemical restrictions on the ionic strength of the buffer are non-ideal conditions for cells.
Buican et al.9 first proposed the use of optical forces for the deflection of particles through a fluidic channel. The force exerted on a particle by an optical beam is a function of the optical power and the relative optical properties of the particle and its surrounding fluid medium. Forces on the order of 1 pN/mW can be achieved for biological cells approximately 10 μm in diameter. While the optical force is small, the force necessary to deflect a cell into an adjacent flowstream is also small, e.g. 900 pN to move a 10 μm diameter cell, 20-40 μm laterally across the flow in a few milliseconds. This is the force necessary to overcome the viscous drag force on the cell at the velocity implied by this lateral motion.
The principles behind the optical forces and general background technology may be found in U.S. Pat. No. 6,744,038, which is incorporated herein by reference as if fully set forth herein.